Method and apparatus for high-gain magnetic resonance imaging

ABSTRACT

A method and apparatus for Magnetic Resonance Imaging with specialized imaging coils possessing high Signal-to-Noise-Ratio (SNR). Imaging and/or Radio Frequency receiving coils include a ballistic electrical conductor such as carbon nanotubes, the ballistic electrical conductor having a resistance that does not increase significantly with length. Due to their enhanced SNR properties, system designs with smaller static magnetic field strength can be constructed for the same quality of imaging, leading to substantial reductions in system size and cost, as well as to enhanced imaging with existing MRI systems.

CROSS-REFERENCES TO RELATED APPLICATIONS

This application is a continuation entitled to the benefit of U.S.patent application Ser. No. 11/582,094 entitled “METHOD AND APPARATUSFOR HIGH GAIN MAGNETIC RESONANCE IMAGING”, filed Oct. 17, 2006, hereinincorporated by reference.

U.S. patent application Ser. No. 11/582,094 is entitled to the benefitof U.S. Provisional Patent Application No. 60/727,924, entitled “METHODAND APPARATUS FOR HIGH GAIN MAGNETIC RESONANCE IMAGING”, filed Oct. 18,2005, herein incorporated by reference.

FIELD OF THE INVENTION

This invention relates to diagnostic medical imaging, and moreparticularly to Magnetic Resonance Imaging at highSignal-to-Noise-Ratios.

BACKGROUND OF THE INVENTION

Magnetic Resonance Imaging (MRI) technology is common today in largermedical institutions worldwide, and has led to huge benefits in thepractice of medicine. A significant factor impeding further increaseduse of this versatile imaging technology is the typically high cost ofboth purchase and maintenance of MRI systems.

The costs associated with the design and manufacture of such systems aredue mainly to the need to generate large and very homogeneous staticmagnetic fields, as well as the need to generate gradient fields forimaging with such systems. Such large static fields are currentlyrequired to obtain high image quality and resolution.

SUMMARY OF THE INVENTION

The present invention includes a method and apparatus for MagneticResonance Imaging that uses imaging coils with highSignal-to-Noise-Ratio (SNR) to tradeoff increased SNR for lower staticmagnetic field strength. Various coil arrangements and systemconfigurations are, disclosed that exploit the electrical properties oflow-resistance imaging coils made of electrical conductors (called“ballistic conductors”) having resistance that does not increasesignificantly with length of conductor, such as electrical conductorsmade from carbon nanotube material. An imaging system of the inventioncan include a homogeneous static magnetic field, or a specially-tailoredinhomogeneous static magnetic field. The present invention also providesa method for constructing high-quality integrated imaging systems thatare also low in weight, and in some cases, sufficiently low in weightand bulk so as to be portable.

The imaging coils of the invention are used to acquire radio frequency(RF) signals emitted by precessing proton spins in a sample to beimaged. Further, similar coils can also be used as transmitter coils totransmit RF signals needed to flip the proton spins in the sample,thereby causing them to precess.

The configuration of the imaging coils of the invention exploits theballistic conductance property of the carbon nanotubes so as to optimizethe magnetic field B1 within an imaging volume produced by the coil perunit current flowing through the coil. According to the invention,carbon nanotube conductors are made into thin films which are thenformed into coils and coil arrangements that create a relatively largemagnetic field B1 within the imaging volume, while also havingrelatively small electrical resistance, thereby providing very high SNRvalues.

Further, the ballistic conductance of the carbon nanotube material ofthe coils enables creation of RF receiving circuitry with large QualityFactors. Large Quality Factors imply maximum power reception, therebyproviding high signal quality and narrow bandwidth so as to provide aMagnetic Resonance Imaging system having highly selective imagingcapabilities. The narrow bandwidth that can be thereby attained can beused for efficient imaging in the presence of an inhomogeneous staticmagnetic field, where slice selectivity for imaging is thereby enhanced.

Due to the significantly enhanced SNR properties of the receiving coilsof the invention, a smaller static magnetic field strength B1 can beused to provide the same quality of imaging as is possible with themagnetic field strength B1 of known magnet systems. Thus, a smallermagnet system can be used in the invention, thereby resulting in amagnetic resonance imaging system that is of significantly reduced size,and of significantly reduced system cost. Conversely, a standard staticfield magnitude B1 and the coils and coil configuration of the inventioncan be used to obtain images of significantly improved image quality andimage resolution as compared to those obtained with standard imagingcoils.

For example, the Magnetic Resonance Imaging system of the invention,receiving electrical signals from at least one RF receiving coilincluding an electrical conductor consisting essentially of carbonnanotube material, can acquire real-time three dimensional volumetricimage data of volume at least 7 cm×7 cm×7 cm, with a voxel size lessthan or equal to 1 mm×1 mm×1 mm, and at a rate of better than 10 framesa second. In one general aspect of the invention, a magnetic resonanceimaging system is provided having an enhanced signal-to-noise ratio. Thesystem includes a magnet system for providing a static magnetic fieldhaving a low magnetic field strength; and at least one RF receiving coilassembly including a coiled electrical conducting element consistingessentially of carbon nanotube material; and RF receiving circuitryhaving a large quality factor.

In a preferred embodiment, the quality factor is substantially 100. Inan alternate embodiment, the quality factor is greater than 15.

In another preferred embodiment, the magnet system for providing astatic magnetic field provides a magnetic field of strength of less thanabout 1.5 Tesla. In an alternate embodiment, the magnet system forproviding a static magnetic field provides a magnetic field of strengthof between 0.1 to 1.5 Tesla. In other embodiments, the magnet system forproviding a static magnetic field provides a magnetic field of strengthof between 0.02 Tesla and 3.00 Tesla.

In some preferred embodiments, a plurality of RF receiving coils isincluded, each RF receiving coil including a coiled electricalconducting element consisting essentially of carbon nanotube material,and the plurality of RF receiving coils is configured as a phased arrayso as to enhance signal acquisition.

In a preferred embodiment, the at least one RF receiving coil hasbetween 3 and 500 complete windings of electrical conducting elementconsisting essentially of carbon nanotube material. In otherembodiments, the at least one RF receiving coil has at least 3 completewindings of electrical conducting element consisting essentially ofcarbon nanotube material. In a further embodiment, the windings ofelectrical conducting are of a winding diameter between 5 cm to 80 cm.In another further embodiment, the windings of electrical conducting areof a winding thickness between 20 nanometers and 500 microns.

In yet other preferred embodiments, the at least one RF receiving coilassembly includes an opposed pair of coiled electrical conductingelements, each coiled electrical conducting element consistingessentially of carbon nanotube material in the form of a thin carbonnanotube film structured as a tightly wound ring of carbon nanotube filmof a film thickness between 20 nanometers and 500 microns, and at least3 turns of winding.

In still other preferred embodiments, the at least one RF receiving coilincludes multiple layers of windings of electrical conducting elementconsisting essentially of carbon nanotube material.

In preferred embodiments, the magnet system is a permanent magnet.

In other preferred embodiments, the at least one RF receiving coil canbe moved by an operator. In alternate preferred embodiments, the atleast one RF receiving coil includes a sensor to sense position andorientation of the at least one RF receiving coil.

In yet other preferred embodiments, the RF receiving coil includes acoiled electrical conducting element having a plurality of winding turnsin a spiral pattern.

In still other preferred embodiments, the RF receiving coil includes acoiled electrical conducting element having a plurality of winding turnsin a conical helical pattern. In a further preferred embodiment, thecoiled electrical conducting element having a plurality of winding turnsin a conical helical pattern is of an axial length L of as much as 9 cm.In another further embodiment, the coiled electrical conducting elementhaving a plurality of winding turns in a conical helical pattern has alargest winding diameter in the range of 5 cm to 80 cm, and a smallestwinding diameter being up to 20 times smaller than the largest windingdiameter. In yet another further embodiment, the coiled electricalconducting element has a plurality of winding turns in a conical helicalpattern, and the plurality of winding turns has between 3 winding turnsto 500 winding turns.

In a preferred embodiment, the RF receiving coil includes a first coiledelectrical conducting element having a plurality of winding turns in aconical. helical pattern followed by at least a second coiled electricalconducting element having a plurality of winding turns in a coniccalhelical pattern

In another preferred embodiment, the RF receiving coil includes a coiledelectrical conducting element having a plurality of winding turns in atightly wound ring pattern.

In yet another preferred embodiment,. the at least one RF receiving coilincludes a coiled electrical conducting element having a plurality ofwinding turns in a spiral pattern. In a further embodiment, the coiledelectrical conducting element has between 3 and 500 complete windings ofcarbon nanotube material. In a yet further embodiment, the innermostwindings of the coiled electrical conducting element are of a windingdiameter of substantially 1 cm, and wherein the outermost windings ofthe coiled electrical conducting element are of a winding diameter ofbetween 5 cm to 80 cm. In a still further embodiment, the electricalconducting element having a plurality of winding turns in a spiralpatter, the plurality including at least three complete turns ofwinding. In another further embodiment, the RF receiving coil furtherincludes a second coiled electrical conducting element having aplurality of winding turns in a spiral pattern, the second coiledelectrical conducting element being serially connected, with the samesense of winding, to the coiled electrical conducting element having asensor to sense position and orientation. In yet another furtherembodiment, the coiled electrical conducting element having a sensor tosense position and orientation and the second coiled electricalconducting element are separated by a distance between 2 mm and 15 mm.In still another further embodiment, further including between three andfifteen additional coiled electrical conducting elements in serialconnection.

In another general aspect of the invention, a magnetic resonance imagingsystem includes a magnet system for providing a static inhomogeneousmagnetic field within an imaging volume, such that the staticinhomogeneous magnetic field is stronger in some portions of the imagingvolume than in other portions of the imaging volume where the staticmagnetic field is weaker; and at least one RF receiving coil including acoiled electrical conducting element consisting essentially of carbonnanotubes, the at least one RF receiving coil being of a configuration,and being positioned, such that: a magnetic field of the RF receivingcoil is stronger in some portions of the imaging volume than in otherportions of the imaging volume where the magnetic field of the RFreceiving coil is weaker, and the magnetic field of the RF receivingcoil is stronger in the portions of the imaging volume where the staticinhomogeneous magnetic field is weaker.

In another general aspect of the invention for use in a magneticresonance imaging system, an RF receiving coil assembly is provided thatincludes a coiled electrical conducting element consisting essentiallyof carbon nanotube material; and RF receiving circuitry having a largequality factor.

In still another general aspect of the invention, a magnetic resonanceimaging system is provided having an enhanced signal-to-noise ratio. Thesystem includes a magnet system for providing a magnetic field; and atleast one RF receiving coil assembly including: a coiled electricalconducting element consisting essentially of carbon nanotube material;and RF receiving circuitry having a quality factor greater than 15.

BRIEF DESCRIPTION OF THE DRAWINGS

The invention will be more fully understood by reference to the detaileddescription, in conjunction with the following figures, wherein;

FIG. 1 is an illustration of a magnet system and an associatedinhomogeneous static magnetic field pattern, as indicated by a pluralityof constant magnitude iso-surfaces;

FIG. 2 is an illustration of a ballistic conductor imaging coil in theform of a tightly wound ring coil;

FIG. 3 is an illustration of a ballistic conductor imaging coilconfiguration in the form of a pair of phased array ring coils;

FIG. 4 is an illustration of a ballistic conductor imaging coil in theform of a generalized helical coil;

FIG. 5 is an illustration of a ballistic conductor imaging coilconfiguration in the form of a repeated generalized helical coil;

FIG. 6 is an illustration of a ballistic conductor imaging coil in theform of a spiral pancake coil;

FIG. 7 is an illustration of a ballistic conductor imaging coilconfiguration in the form of a layered spiral pancake coil;

FIG. 8 is an illustration of a ballistic conductor imaging coilconfiguration in the form of a trio coil;

FIG. 9 is an illustration of a ballistic conductor imaging coilconfiguration in the form of an angled pair coil;

FIG. 10 is an illustration depicting three positions of a moveableballistic conductor imaging coil configuration having a coil in each ofthree locations and their respective imaged volumes, as used in dynamicscanning of a complete three dimensional volume of interest;

FIG. 11 is a schematic illustration of a first magnet system embodiment;

FIG. 12 is a schematic illustration of a second magnet systemembodiment; and

FIG. 13 is a schematic illustration of a third magnet system embodiment

DETAILED DESCRIPTION

The imaging system of the present invention includes a magnet systemhaving at least one magnet for producing either a homogeneous or aninhomogeneous static magnetic field within the desired volume to beimaged, together with at least one Radio Frequency (RF) transmissioncoil, and at least one RF receiving coil made of a ballistic electricalconductor, such as carbon in the form of carbon nanotubes.

Carbon nanotubes have many interesting electrical, mechanical andthermal properties. Specifically, they possess the property of ballisticelectron transport, wherein the electrons transported by the conductordo not get significantly scattered during transport, such that theelectrical resistance offered by the conductor to a current does notincrease with length. In contrast, the resistance of a standard(metallic) electrical conductor increases linearly with length, otherthings being equal. Furthermore, ballistic conductors do not exhibit askin effect such that resistance increases with frequency; in fact inthe MHz frequency range characteristic of MR Imaging, carbon nanotubesdemonstrate a weak decreasing dependence of resistance on frequency (forinstance, this is discussed in Y. P. Zhao et al, Physical Review B,Volume 64, 2001, p. 201402(R)).

Recently, a process has been developed to fabricate useful lengths ofcarbon nanotube conductors in the form of thin sheets (M. Zhang et al,Science, Aug. 19, 2005, p. 1215). These thin sheets can be as thin as 50nanometers. A receiving coil constructed of such a sheet can have verylow intrinsic resistance due to its ballistic conductance properties andthe absence of the skin effect common to metallic (scattering)conductors.

Image quality and image resolution in MRI depends directly on theSignal-to-Noise-Ratio obtained with the imaging coils used with thesystem. The overall Signal-to-Noise-Ratio (SNR) in Magnetic ResonanceImaging depends on the static field strength used in the system, as wellas properties of the imaging coil used, and is determined by thefollowing proportionality relationship:

SNR˜(B ₀)² B ₁ /sqrt(R _(eff))   (1)

where B₀ is the static field strength, B₁ is the field produced by theimaging coil per unit current flowing through it, andR_(eff)=(R_(body)+R_(coil)) is the effective resistance (also called netresistance) due to resistive losses in the patient body and in theimaging coil itself.

Thus, a more sensitive imaging coil (i.e., one with a larger B₁ fieldand a smaller effective resistance R_(eff)) can yield the same SNR whenused with a smaller static magnetic field strength B₀; from the aboverelationship it is evident that every factor of 4 increase in coil SNRvalue, B₁/sqrt(R_(eff)), gives a twofold decrease in static fieldstrength B₀ required to maintain the same SNR value. Correspondingly,the size of the MRI system needed for given image quality or resolutiondecreases as the coil SNR increases.

Imaging coils made of a ballistic conductor, such as carbon in the formof carbon nanotubes, can offer an order of magnitude increase ineffective SNR. This is due to both the decreased intrinsic resistanceR_(coil) and the larger B₁ fields that can be produced: by many windingturns of such a conductor (the resistance of this conductor does notincrease significantly with increased length of winding) that can beused in the imaging coil. Fabricated in the form of a thin film, theweight of the imaging coil can be made small as well, even with manywinding turns.

While the coils of the invention are made of carbon nanotube conductors,it is known in the art to connect the coils to electronic circuitry thattypically employs metallic conductors, semiconductors; or a combinationof metallic and semiconducting components. For example, a known methodfor connecting a carbon nanotube coil to a metallic element or electrodeis discussed in Z. Wu et al, Science, Vol. 305, 27 August 2004, p. 1273.Thus, one of average skill in the art can fabricate a carbon nanotubecoil of the invention in a-form suitable for use with a MagneticResonance Imaging system by simply winding the carbon nanotube conductorin a desired geometry on a suitable substrate, such as a polymeric film,and attaching metallic conductors to the free ends of the carbonnanotube conductor. It is known to attach metallic conductors usingwell-known methods of deposition of a conducting metal, such as gold,copper, palladium, platinum, or silver. It is also known that such adeposition can be performed using known methods by direct application ofa molten metal paste to form a carbon nanotube-metal junction, vapordeposition with application of mechanical pressure, using a sputteringprocess, or similar various other means familiar to those skilled in theart. As described in the above-mentioned reference, standardmicrolithographic techniques or masked deposition techniques are usefulin this regard.

As is known in the art, such as taught in M. Zhang et al, Science, Aug.19, 2005, p. 1215, carbon nanotube sheets can be attached serially toone another by taking advantage of the microwave absorption propertiesof carbon nanotubes. Under suitable irradiation with microwaves causinglocalized melting, nanotube nanotube junctions can be formed, as well asnanotube-polymer substrate junctions or sandwiched constructions ofnanotube-nanotube junctions encased in a surrounding substrate. By thesemeans, if so required, distinct nanotube sheets, or stacks of sheets,are fused to one another while maintaining electrical connectivity. Inone embodiment, such an extension technique or method of effectivelycreating a long length of conductor is used for the construction ofspecialized coil geometries, including multiple turns or windings ormore complex constructions, including the geometries described elsewherein the present invention.

It is known that carbon nanotube conductors can be either Single-Walledor Multi-Walled, and methods of construction of both are known anddescribed in the literature While Multi-Walled Nanotubes are employed inthe sheet drawing method described in M. Zhang et al, Science, Aug. 19,2005, p. 1215, for example, K. Hata et al, Science, Vol. 306, 19November 2004, p. 1362, describes a technique for the water-assistedsynthesis of Single-Walled Carbon Nanotubes. This technique can providepatterned, highly organized nanotube structures including sheets andpillars and nanotube forests, from which further macroscopic structuressuch as sheets or films can be fabricated by means of a drawing process.It's known that the growth of the initial nanotube structures or forestscan often benefit from the presence of catalysts, such as Ironnanoparticles, together with a suitable substrate such as Silicon. Insome cases a suitable doping agent, such as Hydrogen, can yield furtherdecreases in resistance of sheets drawn from the nanotube forests. Theadvantages of doping of both Single-Walled Nanotubes and Multi-WalledNanotubes are described, for example, in M. Zhang et al, Science, Aug.19, 2005, p. 1215.

The examples of film or sheet construction methods in the above arediscussed for illustrative purposes only. Those skilled in the art candevise alternate fabrication or construction methods without departingfrom the scope of the invention as claimed.

As an additional benefit, the low resistance of the imaging coils of thepresent invention can be used to build a coil assembly includingreceiving electronics circuitry with large Quality Factors. This meansthat little energy is dissipated in the circuitry, and more of the rawsignal, is available for amplification and subsequent processing. Italso means that very narrow signal bandwidths can be obtained, so thatvery precise imaging slice selection is possible. This is also usefulfor imaging with an inhomogeneous static field, where it is necessary touse a series of carefully tuned Radio Frequencies over a suitablefrequency range to acquire data from a series of suitably shaped spatialslices for subsequent processing and image reconstruction.

In a preferred embodiment, the carbon nanotube imaging coil is connectedto electronic circuitry to select the correct tuning to ensure maximalresponse within a narrow bandwidth centered around a Radio Frequencywhose signal pickup (from the volume to be imaged) is desired, so as toprovide an optimally resonant signal. It is well-known that thecircuitry can also include signal amplification. One skilled in the artknows that the imaging coil generally has a particular set of electricalcharacteristics such as resistance, impedance and capacitance associatedwith it. One skilled in the art also knows that the electronic circuitrygenerally contains standard components such as resistors, capacitors andinductors arranged is such a manner as to obtain the desired resonantresponse centered around the frequency of interest. Various circuitdesigns are possible and the specific design of such electronics candepend on convenience, ease of implementation and optimality, as isfamiliar to those skilled in the art.

It is also known that the circuitry could also include means of detuningthe imaging coil during a transmit phase when other transmit coils areused to transmit the RF signals that generate proton precession in thesample or anatomy to be imaged. Further, it's known that detuningprevents large currents from being induced in the imaging coil duringthe transmit phase. Such detuning can be accomplished by the appropriateuse of PIN diodes or other standard methods known to those skilled inthe art, and the detuning can be driven by an appropriate voltage signalas is standard. It's known that selectively variable capacitors orvaractors can also be used as part of the electronic reception.circuitry. As is known in the art, the circuit can be designed toprovide a suitably high circuit Quality Factor. Typically, the circuitryis designed to match input impedance into the MR imaging system as thesignal from the electronic circuitry is fed to the MR imaging systemthrough suitable means, such as a coaxial cable. Such circuitry isdescribed, for instance, in E Atalar et al, “High ResolutionIntravascular MRI and MRS using a Catheter Receiver Coil”, MagneticResonance in Medicine, Vol. 36, No: 4, pp. 596-605, October 1996.

In the case where the static magnetic field is inhomogeneous andmultiple Radio Frequencies need to be picked up, coil electronics can beprovided by one of skill in the art to achieve optimal resonant responseacross the entire range of frequencies by suitable active tuning of theelectronic components, for instance by voltage-driven capacitors. Thespecific circuit design elements mentioned here are for illustrativepurposes, and other design elements can be used by those skilled in theart while remaining within the scope of the present invention asclaimed. (See, for example, E. Atalar et al, “High ResolutionIntravascular MRI and MRS using a Catheter Receiver Coil”, MagneticResonance in Medicine, Vol 36, No. 4, pp. 596-605, October, 1996.)

In one embodiment, the same RF coil can be used for both transmissionand reception of RF signals. In another embodiment, different RF coilsare used for transmission and for reception. In the case of either ofthese embodiments, the electrical conductor in at least one receivingcoil is made of a ballistic conductor such as carbon nanotubes. Thestatic magnetic field is produced by at least one magnet suitablydisposed about the imaging volume. The at least one magnet may be apermanent magnet constructed of a high-grade magnetic material such asNeodymium-Iron-Boron, or it can be a superconducting magnet with themagnetic field generated by running large currents through thesuperconductor.

In one preferred embodiment, the magnet system is specialty designed toproduce a tailored inhomogeneity pattern of the magnetic field within asuitable volume intended to be used as the imaging volume. In this case,the spatial field gradient distribution is carefully configured. Therange of variation of the field strength values is sub-divided into aseries of field strength increments, each of which corresponding to, andimaged with, a particular Radio Frequency, thereby implementing sliceselection.

RF pulses are applied over the range of relevant frequencies either as aseries of predominantly single frequency pulses, or as pulsesincorporating several relevant frequencies, or as a completely broadbandsignal encompassing the entire frequency range. These RF pulses serve totip or flip the proton spins in the selected slices. Subsequently, thespins precess to align themselves with the local static field, and inthe process, emit RF radiation picked up by the receiving/imaging coils.In one embodiment, while RF signals emitted by spins precessing in aslice corresponding to a first frequency range are picked up by thereceiving coil, an RF pulse over a second frequency range is transmittedby a separate transmitter coil. In this manner, RF signal acquisitionand transmission can be interleaved. At least one set of gradientfield-producing coils can be present and used to spatially encode protondensity as done with standard MR imaging.

In another preferred embodiment, the static magnetic field used with thecarbon nanotube conductor imaging coil is designed to be homogeneous asin a standard MRI system, and the signal acquisition process follows astandard pattern.

FIG. 1 schematically illustrates an inhomogeneous magnetic fieldpattern. The gradient that exists in this pattern can be used forimaging purposes. More specifically, the isosurfaces 125, 127, 129 and131 of constant magnetic field magnitude produced by the magnet 112 areknown. For illustrative purposes only, typical field strengths withinthe volume of interest may range from 0.1 Tesla to 0.5 Tesla. In onepreferred embodiment, the field strength can be as low as 0.03 Tesla inthe portion of the imaging volume where the static field is lowest andas high as 0.7 Tesla in the portion of the imaging volume where thestatic field is highest. For imaging purposes, the field pattern isdivided into zones in steps of approximately 0.008 Tesla. Such zones115, 118 and 121 are shown in FIG. 1. Each zone is sensitive to RFexcitation tuned to a center frequency corresponding to the associatedfield strength. As the spins in each zone are excited by RFtransmission, a ballistic conductor imaging coil is used to pick up thesubsequent relaxation RF signal as the signals precess back intoalignment with the static magnetic field. The ballistic conductor usedin the imaging coil can provide SNR gains of a factor lying at least inthe range 10 to 20 over existing standard MR imaging coils (made ofmetallic conductors). As mentioned earlier, an increase in SNR by afactor of 16 implies a four-fold reduction in static field strengthmagnitude needed to maintain image quality. Thus, a Magnetic ResonanceImaging system with a 0.35 Tesla static magnetic field strengthconstructed according to the teachings of the present invention canoffer an image quality and resolution comparable to that obtained with astandard 1.5 Tesla MRI system.

FIG. 2 shows an imaging coil 140 embodiment constructed with itselectrical conductor in the form of a thin carbon nanotube filmstructured as a tightly wound ring 141 with a film thickness between 20nanometers and 500 microns and with at least 3 turns of winding. Thewinding diameter can be in the range 5 cm to 80 cm. This figure alsoshows the general direction of the static magnetic field B₀ produced bya permanent magnet that is part of the system (not shown). The patienttable 143 is shown for clarity with its head 145. In the configurationshown the coil is mounted on a lateral side of the patient to be imaged.

Alternatively, FIG. 3 shows an opposed pair 148, 150 of tightly woundrings of carbon nanotube film with a film thickness between 20nanometers and 500 microns, and with at least 3 turns of winding, whichcan be used on either lateral side of the patient. This pair ofrings/imaging coils 148, 150 can be used for MR imaging in standardphased array form. The senses of winding 152 and 154 of coils 148, 150,respectively, are opposed and are indicated for each winding. Thegeneral direction of the static magnetic field B₀ is also shown.

FIG. 4 shows an imaging coil embodiment constructed as a generalizedconical helical carbon nanotube winding with turns of progressivelysmaller winding diameter in the portion of the coil closer to thepatient. The axial length L of the winding can be as high as about 8 cm;the largest winding diameter is in the range 5 cm-80 cm and the smallestwinding diameter can be up to 20 times smaller than the largest windingdiameter. The total number of turns of winding can be in the range 3 to500. The coil is mounted on a lateral side of the patient to be imaged.Again, an opposed pair of such generalized conical helical coils placedon lateral sides, of the patient can be used in phased array form tofurther optimize the SNR.

FIG. 5 is an illustration of an imaging coil embodiment constructed as arepeated generalized conical helical carbon nanotube winding with turnsof progressively smaller winding diameter in portions of the coil closerto the patient. The first winding helix 159 is followed by a secondwinding helix 161. The axial length of the winding of each helix can beas long as about 8 cm; the largest winding diameter is in the range 5cm-80 cm and the smallest winding diameter can be up to 20 times smallerthan the largest winding diameter, After the smallest diameter windingturn of the first winding helix 159, the conductor is routed backaxially 160 to near the axial section of the largest diameter windingturn and the generalized conical helical winding pattern is repeated 161with a possible axial offset. The total number of turns of winding ofeach winding helix can be in the range 3 to 500. The number of distinctwinding helixes can be between 2 and 20 wherein two distinct windinghelixes are shown in FIG. 5. The coil is mounted on a lateral side ofthe patient to be imaged. Again, an opposed pair, or other multiplicity,of such generalized conical helical coils placed on lateral sides of thepatient can be used in phased array form to further optimize the SNR.

FIG. 6 shows an imaging coil embodiment constructed in a spiral pancakewinding pattern with carbon nanotube conductor. The innermost windingdiameter is of the order of about 1 cm and the outermost windingdiameter is in the range 5 cm-80 cm. The total number of turns ofwinding can be in the range 3 to 500.

FIG. 7 shows an imaging coil embodiment including two parallel planarcoils 165 and 167 of spiral pancake windings of carbon nanotubeconductor. The two parallel planar coils are serially connected with thesame sense of winding (about the winding axis) in each planar coil, asshown. Each planar coil can have dimensions as associated with a singlespiral pancake mentioned regarding FIG. 6. The distance between theparallel planar coils can be between 2 mm and 15 mm; there can bebetween two and fifteen such planar coils that are serially connected toform the imaging coil.

FIG. 8 shows a preferred embodiment as a trio of imaging coils 170, 173and 176 seen edge-on. Within the imaging region 179 shown, the net B₁field 181 produced by the trio of coils has a strong radial component.The set of three coils 170, 173 and 176 is used in phased array form ina preferred embodiment. In an alternate embodiment they are seriallyconnected. Each of the individual coils can be any one of the typesdescribed herein and can be constructed according to the invention.Generally, the coil trio partially surrounds a patient during imaging.The angles between the three coils as well as the spacing between themand the overall geometric layout are all carefully chosen to optimizethe size and shape of the imaging region and the B₁ field within it; oneskilled in the art would be familiar with the optimization process. Inone preferred embodiment, the entire coil structure is mounted within arigid frame such that the angles between the coils 170, 173 and 176 andthe entire coil configuration cannot change. In an alternate preferredembodiment the coils 170, 173 and 176 are mounted in a flexible framethat permits changes in angles between the coils 170, 173 and 176 whenthe entire configuration is placed around a patient.

FIG. 9 illustrates a preferred embodiment in the form of angled pair ofimaging coils 184, 186 seen edge-on. The angled pair 184, 186 isconfigured such that within the imaging region 188 shown, asignificantly homogeneous and optimized B₁ field 190 is produced by thecoil pair 184 and 186. Each of the individual coils 184, 186 can be anyone of the types described above. The spacing between the coils 184,186, their size and the angle between them are all carefully chosen tooptimize the size and shape of the imaging region and the B₁ fieldwithin it; one skilled in the art would be familiar with theoptimization process. The coils can be serially connected or they can beused as a phased array. In one preferred embodiment the entire coilstructure is mounted within a rigid frame such that the angle betweenthe coils and the entire coil configuration cannot change. In analternate preferred embodiment the coils are mounted in a flexible framethat permits changes in the angle between the coils when the entireconfiguration is placed around a patient.

In one set of preferred embodiments the above coil configurations canall be used as fixed configurations and immovably located for imagingpurposes. In an alternate set of preferred embodiments each coil typecan be moved around the patient by an operator and placed as convenientnear different sides of the patient to image different regions of thepatient anatomy so as to provide maximum SNR within each region during aprocess of dynamic scanning. In a preferred embodiment of thisinvention, such a movable coil as described here is further equippedwith at least one Micro Electro Mechanical System (MEMS) sensor that canmeasure linear as well as angular accelerations and thence, byappropriate integration, yield spatial positional and orientationalinformation for the imaging coil. This spatial positional informationcan be used by the system to automatically select an optimal and limitedspatial region within which the image is reconstructed as the coil ismoved around the patient in various locations. Thus as the coil is movedaround, image acquisition is tailored to provide a set of optimalimaging regions that are subsequently “stitched” together by systemsoftware to produce a complete reconstruction of the desired patientanatomy.

FIG. 10 depicts dynamic scanning with a single coil that is moved by anoperator and located in the three positions 193, 195 and 197 shown toprovide an optimal image in each of three possibly partially overlappingregions 202, 204 and 206 within the patient cross-section 200. Thesystem software stitches these three regions into a single encompassingregion 208 that provides a complete three dimensional image of theentire volume of interest.

The static magnetic field B₀ can be designed to be homogeneous as in astandard MRI system, or it can be designed to produce a specific, known,magnetic field pattern with known gradients everywhere in the region ofinterest. In the latter case, the tuning of the imaging coil andassociated circuitry is designed to permit variable tuning over a rangeof resonant frequencies that corresponds to the range of static fieldstrengths within the imaging volume. For example a varactor can be usedto achieve such variable tuning and switch between different frequencybands. Additionally, the patient table, the system magnet, or theimaging coil, or any combination of these, can be moved and repositionedto scan different portions of anatomy for subsequent three dimensionalvolumetric reconstruction of the entire anatomy of interest. Specializedimage acquisition and processing modalities, as known to those skilledin the art, can be used to acquire real-time or dynamic images ofnon-static organs such as a beating heart. One consequence of usingimaging coils with ballistic conductors as disclosed in the presentinvention is that the higher SNR available with such coils permitsfaster image acquisition, so that dynamic images can be acquired in amuch more real-time fashion than conventionally possible. The number ofgradient coils used for spatial encoding of proton density can vary from1 to 8 in MRI systems built according to the teachings of the presentinvention, as a non-limiting example provided for purposes ofillustration.

Several magnet system configurations are possible which can be used foroptimal imaging together with the imaging coil designs described herein.FIG. 11 shows a preferred embodiment of a permanent magnet configuration210 in a trans-axial view (view from patient feet) with the magnet belowthe patient. Within the imaging volume the B₀ field lines 212 cross theimaging volume laterally from one side to another as shown. In onepreferred embodiment there can be a vertical gradient in the magneticfield. In an alternate preferred embodiment the magnet is carefullydesigned to produce a substantially homogeneous magnetic field withinthe imaging volume. In another preferred embodiment the magnet can belikewise located above the patient. In still another embodiment therecan be more than one permanent magnet (one above and one below thepatient, for instance) and each magnet is designed to optimize net fieldand gradient properties within the imaging volume. The ring coil of FIG.2, the conical helical coil of FIG. 4, the spiral pancake coil of FIG.6, the layered spiral pancake coil of FIG. 7, the repeated helical coilof FIG. 5, and the angled pair coil of FIG. 9, can all be used with thissystem magnet design.

In a second preferred embodiment of system magnet, there are two magnetson either lateral side of the patient. FIG. 12 shows this preferembodiment of a permanent magnet configuration in a trans-axial view(view from patient feet) with the magnets 215 and 218 disposed laterallywith respect to the patient, together with schematic B₀ static fieldline patterns 223. The table edge 220 from patient feet perspective isalso shown for clarity. Any of the coil types previously described canbe used with this system magnet configuration.

FIG. 13 illustrates another preferred embodiment of system magnet, wherethe field lines 235 emanate from the magnet 230 directly into theimaging volume 237. In the trans-axial view shown (view from patientfeet), we can also see the imaging coil 232 edge-on. The imaging coilcan be any one of the following types previously described: ring coil(FIG. 2), helical coil (FIG. 4), spiral pancake coil (FIG. 6), layeredspiral coil (FIG. 7), repeated helical coil (FIG. 5), or the trio coilconfiguration (FIG. 8). A useful property of the system configurationhere is that in this system and coil configuration, while B₀ decreasesin magnitude away from the magnet, the B₁ field produced by the imagingcoil increases away from the magnet. Thus, the effective SNR in thissystem embodiment can be high everywhere within the imaging volume inlight of equation (1). By suitable design and positioning of the imagingcoil(s) together with the system magnet, the SNR within the imagingvolume can be everywhere optimized.

A method of magnetic resonance imaging can also include receivingelectrical signals using at least one RF receiving coil including acoiled electrical conducting element consisting essentially of carbonnanotube material, and including at least one sensor to estimateposition and orientation of the RF receiving coil. This method includesacquiring high signal-to-noise-ratio RF signals from an imaging volumewith the receiving coil in the presence of a static magnetic field withthe receiving coil in a fixed spatial configuration; moving the RFreceiving coil to a second spatial configuration and acquiring RFsignals from the imaging volume with the RF receiving coil in the secondspatial configuration; using the position and orientation information ofthe RF receiving coil in both configurations to process the RF signalsthereby obtained to calculate image intensity in at least two imagesub-volumes; and automatically combining the image intensities in thetwo image sub-volumes so as to create a three dimensional imageintensity reconstruction of the entire imaging volume.

In a preferred embodiment, more than two receiving coil spatialconfigurations are used to reconstruct more than two corresponding imagesub-volumes from which the three dimensional image intensityreconstruction of the entire imaging volume can be automaticallycombined.

In another preferred embodiment, the position and orientation of the RFreceiving coil is also used to normalize the image intensitydistribution of the reconstructed image.

Other system and coil configurations and variations besides the onesdescribed above can be designed by those skilled in the art of MRImaging and following the teachings herein.

Other modifications and implementations will occur to those skilled inthe art without departing from the spirit and the scope of the inventionas claimed. Accordingly, the above description is not intended to limitthe invention except as indicated in the following claims.

1. A magnetic resonance imaging system comprising: a magnet system forproviding a static magnetic field; and at least one RF transmission coilassembly having at least one electrical conducting element consistingessentially of carbon nanotube material.
 2. A magnetic resonance imagingsystem comprising: a magnet system for providing a static magneticfield; at least one RF transmission coil assembly having at least oneelectrical conducting element consisting essentially of carbon nanotubematerial; and at least one RF receiver coil assembly having at least oneelectrical conducting element consisting essentially of carbon nanotubematerial.
 3. The system of claim 2, where the at least one electricalconducting element in the at least one RF transmission coil assembly isthe at least one electrical conducting element in the at least one RFreceiver coil assembly.
 4. The system of claim 2, where the at least oneRF receiver coil assembly includes electronic circuitry for tuning andimpedance matching.
 5. The system of claim 3, where the at least one RFtransmission coil assembly and the at least one receiver coil assemblyshare electronic circuitry for tuning and impedance matching.
 6. The RFtransmission coil of claim 1, wherein the at least one electricalconducting element is capable of carrying electrical current densitiesin excess of current densities that can be supported in a normalmetallic conductor of similar thickness.
 7. For use in a magneticresonance imaging system, an RF transmission coil assembly including: anelectrical conducting element consisting essentially of carbon nanotubematerial that can carry current densities in excess of those that can besupported in a normal metal of similar thickness.